The present invention relates generally to diagnostic imaging systems and, more particularly, to a reflector for a scintillator array having an integrated air gap. Specifically, the scintillator array is constructed such that a uniform air gap or void exists between adjacent scintillators.
Typically, in computed tomography (CT) imaging systems, an x-ray source emits a fan-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.
Generally, the x-ray source and the detector array are rotated about the gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom.
Typically, each scintillator of a scintillator array converts x-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
Scintillator arrays typically incorporate a reflector layer between adjacent scintillators to limit cross-talk between the scintillators thereby improving light collection efficiency of the corresponding photodiodes. Generally, the reflector is formed of a material comprising chromium oxide or other types of optically absorbent material. Because chromium oxide operates as a good absorbent of light, the relative reflectivity of the reflector is reduced. As such, incorporating a reflector layer that includes chromium oxide, a trade-off in CT detector design is made between lower cross talk and reflectivity. If the reflector layer is fabricated without chromium oxide or other optically absorbent materials, cross talk between scintillators increases. In contrast, implementing optically absorbent materials reduces cross talk but lowers the reflectivity of the reflector. Reduced reflectivity degrades low signal performance and increased cross talk affects spatial resolution. Low signal performance is a function of noise generated in the CT detector. As reflectivity falls the light output of the scintillator also falls. Noise, however, is relatively constant, therefore, decreases in light output increases the ratio of noise to functional light output. Additionally, known CT detectors are constructed such that the reflector material is disposed such that it fills any spaces that exist between adjacent scintillators. This also contributes to increased cross-talk between scintillators as there is a constant interface between the scintillators.
It would therefore be desirable to design a CT detector having a reflector with integrated air gaps to improve light collection efficiency and lower cross-talk between scintillators.